Ultrasonic transducers used for medical imaging have numerous characteristics that lead to the production of high quality diagnostic images. Among these are broad bandwidth, affecting resolution and high sensitivity, which combined with pressure output affects depth of field, to low level acoustic signals at ultrasonic frequencies. Conventionally the piezoelectric materials which possess these characteristics have been made of PZT and PVDF materials, with PZT being particularly popular as the material of choice. However, PZT suffers from a number of notable drawbacks. Firstly, the ceramic PZT materials require manufacturing processes including dicing, matching layer bonding, fillers, electroplating and interconnections that are distinctly different and complex and require extensive handling, all of which can result in transducer stack unit yields that are lower than desired. This manufacturing complexity increases the cost of the final transducer probe and puts design limitations on the minimum spacing between the elements as well as the size of the individual elements. Moreover, PZT materials have a poorly matched impedance to water or biological tissue, such that matching layers need to be added to the PZT materials in order to obtain the desired acoustic impedance matching with the medium of interest.
As ultrasound system mainframes have become smaller and dominated by field programmable gate arrays (FPGAs) and software for much of the signal processing functionality, the cost of system mainframes has dropped with the size of the systems. Ultrasound systems are now available in inexpensive portable, desktop and handheld form, for instance for use as ultrasound diagnostic imaging systems or as ultrasound therapeutic systems in which a particular (tissue) anomaly is ablated using high-energy ultrasound pulses. As a result, the cost of the transducer probe is an ever-increasing percentage of the overall cost of the system, an increase which has been accelerated by the advent of higher element-count arrays used for 3D imaging in the case of ultrasound diagnostic imaging systems. The probes used for ultrasound 3D imaging with electronic steering rely on specialized semiconductor devices application-specific integrated circuits (ASICs) which perform microbeam forming for two-dimensional (2D) arrays of transducer elements. Accordingly it is desirable to be able to manufacture transducer arrays with improved yields and at lower cost to facilitate the need for low-cost ultrasound systems, and preferably by manufacturing processes compatible with semiconductor production.
Recent developments have led to the prospect that medical ultrasound transducers can be batch manufactured by semiconductor processes. Desirably these processes should be the same ones used to produce the ASIC circuitry needed by an ultrasound probe such as a CMOS process. These developments have produced micromachined ultrasonic transducers or MUTs, the preferred form being the capacitive MUT (CMUT). CMUT transducers are tiny diaphragm-like devices with electrodes that convert the sound vibration of a received ultrasound signal into a modulated capacitance. For transmission the capacitive charge applied to the electrodes is modulated to vibrate/move the diaphragm of the device and thereby transmit an ultrasound wave. Since these diaphragms are manufactured by semiconductor processes the devices generally have dimensions in the 10-500 micrometer range, with spacing between the individual diaphragms less than a few micrometers. Many such individual CMUTs can be connected together and operated in unison as a single transducer element. For example, four to sixteen CMUTs can be coupled together to function in unison as a single transducer element. A typical 2D transducer array can have 2000-3000 CMUT transducer elements.
The manufacture of CMUT transducer-based ultrasound systems is therefore more cost-effective compared to PZT-based systems. Moreover, due to the materials used in such semiconductor processes, the CMUT transducers exhibit much improved acoustic impedance matching to water and biological tissue, which obviates the need for a matching layer and yields an improved effective bandwidth.
One of the main challenges in developing effective ultrasound systems, and in particular CMUT transducer-based ultrasound systems is to provide systems with excellent image resolution and good depth-of-field in case of an ultrasound diagnostic imaging system. These are conflicting requirements, as higher frequency pulsed ultrasound leads to improved resolution but shorter depth-of-field due to the frequency dependent attenuation of the medium. In order to obtain high resolution in depth, high pressure short pulses are desired which require a large bandwidth. Although in principle CMUT transducers can generate a broad spectrum of frequencies the bandwidth is limited because the frequency at which they operate efficiently depends strongly on the applied static bias voltage over the CMUT.
B.-H. Kim et al., “An Experimental Study on Coded Excitation in CMUT Arrays to Utilize Simultaneous Transmission Multiple-zone Focusing Method with Frequency Divided Sub-band Chirps,” in Proc. IEEE Ultrasonics Symp., 2013, pp. 1428-1431 disclose the transmission of chirped ultrasound pulses with a CMUT array. However, such pulses exhibit relatively narrow effective bandwidths due to the loss of acoustic performance throughout the bandwidth range and as such are of limited use when trying to improve resolution and/or depth-of-field of the imaging data, for which the acoustical performance should be maintained over an as large as possible bandwidth.